Molecules 2014, 19, 11915-11932; doi:10.3390/molecules190811915

molecules ISSN 1420-3049

www.mdpi.com/journal/molecules

Article

Self-Assembled Micelles Composed of Doxorubicin Conjugated Y-Shaped PEG-Poly(glutamic acid)2 Copolymers via Hydrazone Linkers

Bowen Sui 1, Hui Xu 2,*, Jian Jin 1, Jingxin Gou 2, Jingshuo Liu 1, Xing Tang 2, Yu Zhang 2,

Jinghua Xu 3, Hongfeng Zhang 3 and Xiangqun Jin 1,*

1 Department of Pharmaceutics, College of Pharmacy Sciences, Jilin University, Changchun 130021,

Jilin, China 2 Department of Pharmaceutics, Shenyang Pharmaceutical University, Shenyang 110016, China 3 School of Life Science and Biopharmaceutics, Shenyang Pharmaceutical University,

Shenyang 110016, Liaoning, China

* Authors to whom correspondence should be addressed; E-Mails: xuhui-spu@163.com (H.X.);

jinxq@jlu.edu.cn (X.J.); Tel./Fax:+86-24-2398-6356 (H.X.); Tel./Fax: +86-431-8561-9662 (X.J.).

Received: 11 June 2014; in revised form: 18 July 2014 / Accepted: 21 July 2014 /

Published: 11 August 2014

Abstract: In this work, micelles composed of doxorubicin-conjugated Y-shaped

copolymers (YMs) linked via an acid-labile linker were constructed. Y-shaped copolymers

of mPEG-b-poly(glutamate-hydrazone-doxorubicin)2 and linear copolymers of

mPEG-b-poly(glutamate-hydrazone-doxorubicin) were synthesized and characterized.

Particle size, size distribution, morphology, drug loading content (DLC) and drug release of

the micelles were determined. Alterations in size and DLC of the micelles could be achieved

by varying the hydrophobic block lengths. Moreover, at fixed DLCs, YMs showed a smaller

diameter than micelles composed of linear copolymers (LMs). Also, all prepared micelles

showed sustained release behaviors under physiological conditions over 72 h. DOX loaded

in YMs was released more completely, with 30% more drug released in acid. The anti-tumor

efficacy of the micelles against HeLa cells was evaluated by MTT assays, and YMs

exhibited stronger cytotoxic effects than LMs in a dose- and time-dependent manner.

Cellular uptake studied by CLSM indicated that YMs and LMs were readily taken up by

HeLa cells. According to the results of this study, doxorubicin-conjugated Y-shaped

PEG-(polypeptide)2 copolymers showed advantages over linear copolymers, like assembling

into smaller nanoparticles, faster drug release in acid, which may correspond to higher

OPEN ACCESS

Molecules 2014, 19 11916

cellular uptake and enhanced extracellular/intracellular drug release, indicating their

potential in constructing nano-sized drug delivery systems.

Keywords: miktoarm-star copolymer; conjugate; doxorubicin; pH-responsive drug release;

drug delivery

1. Introduction

Benefitting from the advances in polymer science, more and more novel synthetic polymers are being

used in the construction of nano-scale drug delivery systems (DDS) that target tumors by their enhanced

permeation and retention (EPR) effects [1]. Simple alterations in the morphology, size, release profile or

drug-loading content of nano-carriers can be readily achieved by modifying parameters like the

generations of dendrimers [2,3], grafting percentage of graft copolymers [4,5] and the conformation of

the blocks [6,7]. Recently, miktoarm-star copolymers with branches at the junction point of the linear

copolymers have aroused increasing interest [8–11]. These non-linear copolymers exhibited unique

self-assembly behaviors believed to be based on the conformational entropy reduction caused by the

introduction of branches at the junction point [10,12]. The architecture dependent self-assemble of these

polymers affords conventional and non-conventional nano-structures like micelles [8,10,11],

polymersomes [9,13], multicompartment micelles [14] and wormlike micelles [15]. More recently,

polyamino acids were used to build a variety of polymers due to their high biocompatibility. Also,

secondary structures like α-helix and β-sheets adopted by natural or synthetic polyamino acids endowed

them advantages in terms of stability and processability [10]. The application of miktoarm-star

copolymers for drug delivery has been reviewed by Soliman et al. [16]. For example,

PNIPAM-b-poly(undecylenic acid)2 was synthesized by Li et al. [17] to construct prednisone acetate

loaded thermal-responsive micelles, which showed a temperature-dependent sustained release.

However, the micellar stability might be limited due to the flexibility of the all-carbon backbone of the

polymer, as reflected by a higher CAC value, while the secondary structure adopted by poly(amino

acids) reduced chain mobility, thereby increasing the micellar stability to some extent. Another AB2 type

copolymer (mPEG-b-PBLG2) was synthesized by Li et al. [8], and the tamoxifen loaded micelles also

exhibited a sustained release pattern.

For micellar systems composed of AB type copolymers, increasing the length of the hydrophobic

chain is a direct way to increase the DLC and reduce the drug release rate. However, micelles with longer

hydrophobic chain lengths might be less stable due to the reduced hydrophilic surface coverage [8,18,19].

According to the above mentioned self-assembling properties of miktoarm-star copolymers, it seems

like that the hydrophobic chains of miktoarm-star copolymers could be shorter than the AB type

counterparts with similar degrees of polymerization, which means a better stability and a smaller

diameter of the micelles. Likewise, for miktoarm-star copolymers with similar single hydrophobic chain

lengths, more cargo could be loaded into the micelles, so some new traits might be possible by

constructing micelles with these branched polymers.

Doxorubicin (DOX) is a widely used antitumor drug that has been involved in a great deal of research

into the development of DDSs. The reason for its popularity may be due to its unique physical-chemical

Molecules 2014, 19 11917

properties. The salt form of DOX (DOX·HCl) is water-soluble and can be incorporated into liposomes [20]

and polymersomes [21], and it could form aggregates with anionic polymers through electrostatic

interactions [22]. When neutralized, DOX becomes insoluble in water and can be loaded into the

hydrophobic cores of micelles [23]. Moreover, an acid-labile hydrazone linker could be introduced to

achieve a pH-responsive release [24]. Also, for DDSs carrying DOX or other chemotherapeutic agents,

the drug loading content (DLC) is a key parameter that is closely related to the anti-tumor efficacy and

even multidrug resistance induction. This indicates that further increasing drug loading while

maintaining a small particle size is a task that still remains to be accomplished.

In this study, Y-shaped miktoarm-star copolymers of α-methoxypoly(ethylene glycol)-b-poly

(glutamate-hydrazone-DOX)2 (mPEG-b-p(Glu-Hyd-DOX)2) were synthesized (Scheme 1). In addition,

AB type linear copolymers mPEG-b-poly(Glu-Hyd-DOX) were also synthesized. Both mPEG and poly

(glutamic acid) are biocompatible and have been widely used in various studies. DOX was conjugated to

the polymers through an established hydrazone-linker [25] to achieve a stimuli-responsive release and

less leakage into the bloodstream. In order to confirm the above hypotheses, micelles composed of these

Y-shaped and linear copolymers were prepared and their size, DLC, in vitro release and cytotoxicity

were investigated.

2. Results and Discussion

2.1. Synthesis and Characterization of Polymers

The ring-opening polymerization (ROP) of amino acid NCAs is a commonly employed method in the

facile synthesis of polyamino acids with controllable degrees of polymerization and polydispersity

indices (PDI) [7]. Y-shaped miktoarm-star mPEG-(PBLG)2 copolymers with different molecular

weights were synthesized via ring-opening polymerization of γ-benzyl-L-glutamate-N-carboxy

anhydride (BLG-NCA) using α-methoxy-poly(ethylene glycol) with two primary amino groups at the

ω-terminal as a macromolecular initiator [8]. DOX was conjugated to the polymers via an acid-labile

hydrazone linker with hydrazide groups obtained from EAE aminolysis of the PBLG block. The

synthetic route is shown in Scheme 1.

The chemical structures of the products were confirmed by FTIR and 1H-NMR. Figure 1 shows the

FTIR spectra of BLG-NCA, mPEG-(PBLG)2, mPEG-p(Glu-Hyd)2 and mPEG-p(Glu)2-DOX polymers.

The strong absorptions at 1858 cm−1 and 1786 cm−1 in Figure 1a were the characteristic peaks of the

–CO-O-CO- end group of BLG-NCA, which had disappeared completely in the mPEG-(PBLG)2 FTIR

spectrum in Figure 1b. This indicated the complete transformation of BLA-NCA. The disappearance,

together with the emergence of an amide I and amide II bands at 1651 cm−1 and 1549 cm−1 was

due to the formation of amide groups in mPEG-(PBLG)2, and the marked peak at 1107 cm−1 was

attributed to the C-O-C stretching vibration and the peak at 2889 cm−1 proved the incorporation of

mPEG-S-(CH2CH2NH2)2. The disappearance of the stretch vibration of the ester C=O (1734 cm−1) and

the reduced absorbance at 696 cm−1 and 740 cm−1 (characteristic peaks of mono-substituted phenyl

group) as well as the appearance of amide band 1645 cm−1 and 1531 cm−1 all indicated the conjugation

of hydrazide (Figure 1c). The conjugation of DOX was confirmed by the appearance of peaks at

989 cm−1, 1,384 cm−1 and 3,433 cm−1 and the widening of the peak at 1112 cm−1 in Figure 1c‒e.

Molecules 2014, 19 11918

Scheme 1. Synthetic routes of linear or Y-shaped PEG-poly(Glu-Hyd-DOX) polymers and

formation of micelles.

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Figure 1. Fourier transform infrared spectroscopy spectra of: BLG-NCA (a); PEG-(PBLG)2

(b); PEG-poly(Glu-Hyd)2 (c); PEG-poly(Glu)2-DOX polymers (d) and DOX·HCl (e).

The 1H-NMR spectra of the BLG-NCA and polymers in CDCl3 or CF3COOD/DMSO-d6 are shown in

Figure 2. In Figure 2a, the smooth baseline in the spectrum of BLG-NCA confirmed its high purity since

this monomer is relatively sensitive to humidity and the quality of the polymer is dependent on the purity

of the monomers. The NMR spectrum of the mPEG-(PBLG)2 copolymer in Figure 2b exhibited

characteristic peaks at 3.39 ppm (-CH3O-) and 3.66 ppm (-CH2-CH2-O-), which were assigned to the

proton of the MPEG block. Peaks at 5.06 ppm (-CH2-C6H5), 3.95 ppm (-CH-CO-NH-), 2.63 ppm

(-CH2-CH2-COO-), and 2.15-2.29 ppm (-CH2-CH2-COO-) were assigned to the protons of the PBLG

unit. The degree of polymerization was calculated from the signals at δ3.66 ppm and δ5.06 ppm as they

appeared with good reproducibility and little sign of interference. As shown in Figure 2c, approximately

all of the benzyl esters were replaced by hydrazide groups as the signals of the -CH2- group (δ5.06) were

nearly unnoticeable. The NMR spectra of benzyl ester protected polymers and hydrazinolyzed polymers

are shown in Figure S1and Figure S2.

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Figure 2. 1H-NMR spectra of BLG-NCA (a); PEG-(PBLG)2 (b) copolymer in CDCl3 and

PEG-poly(Glu-Hyd)2 (c) polymer in CF3COOD/DMSO-d6.

The molecular weights of the PEG-PBLG or PEG-PBLG2 copolymers calculated by 1H-NMR are

listed in Table 1. The initiators employed showed an initiating efficiency of no less than 0.8, which

means a controllable constitution of polymers could be achieved by varying the feed ratios of the

initiators and monomers. As verified by the GPC in Figure S3, the polymers synthesized displayed

unimodal molecular weight distributions with little impurities, and a constant PDI less than 1.20 was

also obtained (Table 1). The retention time of PEG5k-(PBLG5.7k)2 is 0.33 min shorter (22.98 min vs.

22.65 min) than that of its linear counterpart with a similar molecular weight. This can be ascribed to the

higher molecular volume of the miktoarm star copolymers.

Table 1. Compositions and some characteristic properties of hybrid polypeptide copolymers.

Copolymer a Feed molar ratio

of [I]/[N] b Mn of copolymer c

(Da) DP d

Initiator efficiency e

Polydispersity f (MW/Mn)

PEG5k-PBLG5.7k 1/30 10694 26 0.87 1.07 PEG5k-PBLG13.6k 1/70 18578 62 0.89 1.19

PEG5k-(PBLG5.7k)2 1/65 16388 52 0.80 1.11 PEG5k-(PBLG6.8k)2 1/75 18578 62 0.83 1.11

Notes: a The subscripts indicate the mean Mn of each block; b [I]/[N] refers to the molar ratio of

initiator (MPEG-NH2 or MPEG-S-(CH2CH2NH2)2) to monomer (BLG-NCA); c determined by 1H-NMR analysis; d refers to the degree of polymerization and calculated by 1H-NMR; e calculated by

DPcalculated/DPtargeted; f measured by GPC.

Molecules 2014, 19 11921

The anti-dilution ability which is usually presented as the CAC of polymeric micelles is a key

parameter that determines the micellar performance in vivo. The CAC of the drug-conjugated

copolymers was characterized by steady-state fluorescent probe studies. Pyrene was chosen as the

fluorescent probe because of its photophysical characteristics and other properties [26]. Pyrene will

preferentially partition into hydrophobic microdomains with a concurrent change in the photophysical

properties of the molecule. The maximum excitation wavelength of pyrene increased from 336 nm to

338 nm versus the polymer concentration as shown in Figure 3. And we found DOX showed extremely

low excitation intensity in the range of 300 nm to 350 nm. Therefore, the fluorescence intensity ratio at

338 and 336 nm (I338/I336) of pyrene versus the logarithm of the polymer concentration was plotted from

the excitation spectra (Figure 3 inset). The CAC of the polymer was calculated to be the corresponding

concentration at the inflection point of the plot representing the formation of micelles. The calculated

CACs of the copolymers are shown in Table 2 and a slightly lower CAC of the Y-shaped copolymers

was observed. This indicated a better thermodynamic stability of YMs compared with LMs. The trend of

reduced CAC values with increased hydrophobic block lengths is in accordance with other’s data [7].

However, we noticed that the CAC values of the Y-shaped copolymers differed from those of other

studies. Both Yang’s [27] and Maglio’s [28] groups compared the CAC values between linear and

Y-shaped copolymers with similar molecular weights, and a higher CAC of the Y-shaped copolymers

was found, which is not in accordance with our results. This might be caused by differences in the

physico-chemical properties of the hydrophobic blocks. Poly(solketal acrlate) and poly(ε-caprolactone)

were employed as hydrophobic blocks in studies by Nie and Quaglia, respectively, and the Y-shaped

copolymers in this CAC test are mPEG-PG2Ds, so the discrepancy in CAC might be related to the

secondary structures adopted by PBLGs and another study involving the same hydrophobic blocks

showed identical trends in CAC changes [29].

Figure 3. The representative plots of I338/I336 versus log C of PEG-p(Glu)2-DOX polymer.

The excitation spectra of polymer solution with different concentrations in the presence of

pyrene are shown in the inset.

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Table 2. Characteristics of different doxorubicin-conjugated micelles.

Sample Polymer Size (nm) Polydispersity index DLC (%) CAC (mg/L)

PEG-PGDI PEG-P(Glu)26-DOX 149.9 ± 3.7 0.081 ± 0.023 9.92 ± 0.25 9.0 PEG-PGDII PEG-P(Glu)62-DOX 231.4 ± 8.3 0.229 ± 0.061 18.8 ± 0.18 4.2 PEG-PG2DI PEG-P(Glu26)2-DOX 141.3 ± 5.2 0.222 ± 0.042 16.2 ± 0.12 5.4 PEG-PG2DII PEG-P(Glu31)2-DOX 165.6 ± 2.4 0.116 ± 0.035 18.2 ± 0.45 3.8

Notes: Data represent mean ± standard deviation (n = 3).

2.2. Preparation of DOX-Conjugated Polymer Micelles

DOX-conjugated micelles were prepared by dialysis. PBS (pH 7.4, 10 mM) and deionized (DI) water

were used as the dispersion medium. In this manner, DOX-conjugated micelles based on PEG hybrid

polypeptides were prepared. The particle sizes, size distributions and DLCs of all prepared micelles are

shown in Table 2. According to the results, the diameter data further confirmed that the particle sizes are

determined by the lengths of the hydrophobic blocks. The morphologies of the prepared micelles were

examined by TEM (Figure 4c,d) and the micellar core-shell structure was clearly presented.

Figure 4. Particle size characterization of prepared micelles. The dynamic laser light

scattering analysis of PEG-PGD and Y-shaped PEG-PG2D micelles (a) or (b), Transmission

electron microscopy images of PEG-PG2DII (c) and PEG-PGDII (d) micelles.

Also, when judged in combination with DLC, the advantage of YMs was clear: for micelles with

similar diameters (PEG-PGDI and PEG-PG2DI in Figure 4(a)), the DLC of PEG-PG2DI is nearly twice as

that of PEG-PGDI (16.2% vs. 9.92%) as the Y-shaped copolymers provided more sites for DOX

Molecules 2014, 19 11923

conjugation at poly glutamic acid block lengths similar to that of the linear copolymers. However, due to

the steric hindrance caused by the loaded DOX, a doubled DLC was not achieved. On the other hand, for

micelles with similar DLCs (PEG-PGDII and PEG-PG2DII), the diameters of the YMs were significantly

smaller (231.4 nm vs. 165.6 nm, Figure 4b) which indicated better accumulation could be achieved as

smaller micelles could more readily pass through tumor vessels [1] and penetrate into tumor tissues [30].

This provided a method to use the limited micellar core in a more efficient manner. The determined

results for the particle size and DLC are in accordance with our previously stated hypothesis.

2.3. In Vitro Drug Release Studies

The drug release behavior of the DOX-loaded LMs and YMs was investigated by dialysis diffusion at

pH 5.0 and 7.4 [31]. The UV absorption spectra of DOX at different pH values (5.0, 7.4) were

determined by UV spectrophotometry. Figure 5a shows DOX release patterns for PEG-PGDI,

PEG-PGDII, PEG-PG2DI, and PEG-PG2DII.

Figure 5. In vitro drug release and cytotoxicity assays. (a) In vitro DOX release profiles of

the DOX-loaded polymer micelles at two different pH values (pH 5.0 and 7.4), Cytotoxicity

of different concentrations of (b) empty polymer micelles and (c) Free DOX or DOX-loaded

polymer micelles against HeLa cells by MTT assay, (d) Cytotoxicity of DOX-loaded

polymer micelles against HeLa cells with different incubation times at the same drug

concentration of 10 μg/mL.

Note: * p < 0.01.

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A pH-dependent release pattern was observed in LMs and YMs with a 10%–26% higher DOX release

at pH 5.0. The enhanced release was caused by acid-dependent breakage of hydrazone linkers [32]. All

micelles demonstrated a similar release pattern at pH 7.4 with less than 45% DOX being released in 72 h.

However, accelerated DOX release occurred at pH 5.0, and YMs showed a significantly higher release

with almost 70% of the encapsulated drug being released in 72 h, which was 15% more than that of the

LMs. The drug release pattern usually depends on a variety of factors like crystallinity, degradation rate,

drug binding affinity, molecular weight, polymer composition, particle size, rate of hydration of the

materials, surface characteristics and swelling ratio [33]. For micelles composed of drug-polymer

conjugate, there are fewer chances for drugs to deposit on the surface of the micelles, which means the

drug release rate of these micelles is primarily determined by the diffusion of the cleaved drug through

the polymer matrix. This indicates that the hydrophilic blocks could to some extent control the release of

the cargo as these hydrophilic blocks covering the micellar core functions like a retarding layer. For

YMs, the branched drug-loading blocks that have larger volume increased the gap between the

shell-forming PEG chains, decreased the density of PEG coverage, weakened the retarding layer thus

facilitated drug release of YMs (Figure 5a). Moreover, to further explain the differences in release rates,

we tested the solubilizing behavior of mPEG-poly(Glu-Hyd) and mPEG-poly(Glu-Hyd)2 with similar

molecular weights in acidic media and found that mPEG-poly(Glu-Hyd)2 was more soluble (at pH 5.0)

than mPEG-poly(Glu-Hyd). So, the differences in release rates could be explained by the relative

hydrophilicity in acidic media of mPEG-poly(Glu-Hyd)2 after breakage of the hydrazone linker and the

shorter chain lengths of the Y-shaped copolymers means less chain entanglement which accelerated the

deformation of micelles. Compared with the established nanomedicine Doxil®, the formulation we

prepared showed a more obvious acceleration in release rate under acidic environment, while pH has

little influence on the release profile of Doxil® [34]. However, Doxil® presented better performance in

sustained release [35]. These results indicate that YMs are able to release their cargoes promptly at the

target site.

2.4. Cell Cytotoxicity Studies

As a drug carrier, biocompatibility is a primary concern, so the cytotoxicity of empty PEG-PGDI and

PEG-PG2DI micelles was studied by MTT assay using HeLa cells. Figure 5b shows the cell viability

after incubation for 24 h with polymeric micelles of empty PEG-PGDI and PEG-PG2DI at different

concentrations. The results obtained showed that the cell viability of HeLa cells decreased as the

concentration of micelles increased and the cell viability assay indicated that the empty PEG-PG2DI

copolymer was generally not very cytotoxic to HeLa cells at a concentration up to 50 μg/mL,

demonstrating the excellent biocompatibility.

To investigate the antitumor activity of DOX-conjugated polymeric micelles, polymeric micelles

incorporating doxorubicin were also examined using HeLa cells. The cell viability after incubation for

72 h with different concentrations of free DOX and DOX-conjugated micelles is shown in Figure 5c. In

general, the viability of cells incubated with the two micelles decreased in a dose-dependent manner.

Moreover, a lower cell viability (p < 0.01) of PEG-PG2DI micelles compared with PEG-PGDI micelles

was observed at all concentrations. The IC50 value of free DOX, PEG-PG2DI micelles and PEG-PGDI

micelles was 0.063, 0.517 and 0.673 μM, respectively. The time-dependent cytotoxicity of PEG-PGDI

Molecules 2014, 19 11925

and PEG-PG2DI micelles in HeLa cells after incubation for 24 h, 48 h and 72 h with the same

concentration (10 μM calculated by DOX) was compared and the results are summarized in Figure 5d.

The cell viabilities decreased gradually as the incubation time increased and, at each time point, the

PEG-PG2DI micelles were more cytotoxic than PEG-PGDI micelles. The mild acidic environments in

tumor tissue [36] and in the endosomal lumen [37] can be used to facilitate drug release from the

synthesized polymer-drug conjugates, and DOX may be taken up by cells via two different routes: part

of the loaded DOX is released in tumor tissue and diffused into tumor cells, and the rest of the drug

residing in micelles are endocyted and liberated within the endosomes. Since the cell viability is closely

related to the amount of liberated DOX, the differences in cell viabilities could be attributed to the faster

DOX release rates of PEG-PG2D micelles.

2.5. Cellular Uptake Study

Figure 6. Uptake of DOX-loaded micelles by HeLa cells. Confocal laser microscopy

(CLSM) images of HeLa cells incubated with DOX·HCl (a), DOX-conjugated PEG-PGDI (b)

and PEG-PG2DI (c) micelle for 4 h with a DOX concentration of 10 μg/mL (scale bar 50 μm).

Molecules 2014, 19 11926

Confocal laser microscopy (CLSM) was used to investigate the cellular uptake and intracellular

distribution of free DOX, PEG-PGD and PEG-PG2D micelles in HeLa cells after a 4 h incubation. As

shown in Figure 6, red fluorescence was observed in the cytosol, which indicated that the PEG-PG2D

micelles and PEG-PGD micelles could be readily internalized by HeLa cells (Figure 6b,c). The

relatively stronger fluorescence of PEG-PG2D micelles was caused by the accelerated DOX release after

being internalized in the acidic lumen endosomes that was in accordance with the results of the drug

release study. In Figure 6a, strong fluorescence of DOX was observed, which was due to the different

cellular uptake mechanisms of free DOX and micelles. The free DOX could pass through the cell

membrane by diffusion which is driven by a concentration gradient across the membrane, while micelles

were taken up by endocytosis which is both time- and energy-consuming [38,39]. Since DOX is a

compound that is highly permeable to cellular membrane, the difference between YMs and LMs in

intracellular drug delivery was not clearly presented, which means a further study employing agents with

poor cellular membrane permeability is warranted. According to the CLSM results, after entering the

cell, free DOX can diffuse rapidly into the nuclei where they exert cytotoxic effects by inhibiting DNA

replication [24]. However, DOX was liberated from the PEG-PGD or PEG-PG2D micelles conjugated

with DOX in a controlled manner in the acidic endosomes, and then diffused into the cytosol before it

finally entered the cell nuclei.

3. Experimental Section

3.1. Materials

α-Methoxyl-ω-amino poly(ethylene glycol) (mPEG-NH2, Mn: 5,000, polydispersity index: 1.02,

Seebio Biotech, Inc. Co., Shanghai, P.R. China) and α-methoxyl-ω-N,N-bis(aminoethyl) poly(ethylene

glycol) (mPEG-S-(CH2CH2NH2)2, Mn: 5,000, polydispersity index: 1.10, Seebio Biotech, Inc. Co.)

were dried by azeotropic distillation with toluene; γ-benzyl-L-glutamate (BLG, Sichuan Tongsheng

Amino Acid Ltd. Co., Deyang, China) and triphosgene (99%, GL Biochem, Shanghai, China) were dried

before use; doxorubicin hydrochloride (DOX·HCL, Beijing Huafeng United Technology Ltd. Co.,

Beijing, China) was used as received. Triethylamine, hydrazine hydrate, anhydrous ethyl ether,

methanol, dimethylsulfoxide (DMSO) and all other organic solvents used in the synthesis experiments

were purchased from National Medicine Chemical Reagent Ltd. Co. (Shanghai, China). Other solvents

were all dried before use. The water used in all experiments was double distilled. RPMI 1640 cell culture

medium, fetal bovine serum (FBS), penicillin-streptomycin, trypsin, Dulbecco’s phosphate buffered

saline (DPBS), dimethyl sulfoxide (DMSO) and 3-(4,5-dimethylthiazol-2-yl)-2,5-diphenyltetrazolium

bromide (MTT) were purchased from Gibco Company (New York, NY, USA).

3.2. Synthesis of γ-benzyl-L-glutamate-N-carboxy anhydride

γ-Benzyl-L-glutamate-N-carboxy anhydride (BLG-NCA) was prepared according to the published

method [40] with some modifications. Briefly, BLG (8.0 g, 54.4 mmol) was suspended in dry THF

(80 mL) and heated to 55 °C. After 30 min, a solution of triphosgene (4.0 g, 13.5 mmol) in anhydrous

THF (40 mL) was added dropwise under stirring, and the reaction was conducted under argon

atmosphere until the mixture became transparent (about 2 h). The solvent was removed by rotary

Molecules 2014, 19 11927

evaporation under reduced pressure and the crude product was washed with anhydrous hexane to

remove residual triphosgene. Then, the crude crystals of BLG-NCA were recrystallized three times in a

mixture of EtAc, THF and petroleum ether at room temperature to obtain white crystalline needles.

3.3. Synthesis of Methoxyl Poly (ethylene glycol)-Poly (γ-benzyl-L-glutamate) Diblock and AB2

Y-Shaped 3-Miktoarm Star-Block Copolymers

mPEG-Poly(γ-benzyl-L-glutamate)(PBLG) diblock and mPEG-(PBLG)2 3-miktoarm star copolymer

were synthesized by ring-opening polymerization (ROP) of BLG-NCA initiated by mPEG-NH2 or

mPEG-S-(CH2CH2NH2)2 in anhydrous DCM with different feed ratios. The polymerization reaction was

carried out with stirring under argon atmosphere at 25 °C for 72 h. The mixture was evaporated under

reduced pressure to reduce the volume, and then precipitated twice with anhydrous ethyl ether. The

precipitate was collected by filtration, dissolved in THF and dialyzed (cellulose ester, MWCO: 3500)

against deionized water. Finally, the fluffy white powder was collected by freeze-drying.

3.4. Synthesis of Methoxyl Poly (ethylene glycol)-Poly (glutamate-hydrazone-doxorubicin)2 Polymer

The benzyl esters of mPEG-PBLG diblock copolymer and mPEG-(PBLG)2 miktoarm-star copolymer

were substituted with hydrazide (Hyd) by an ester-amide exchange (EAE) aminolysis reaction to obtain

mPEG-poly(Glu-Hyd) and mPEG-poly(Glu-Hyd)2 polymers [24]. Excess hydrazine hydrate (20 times

the number of BLG repeating units) was reacted with mPEG-PBLG or mPEG-(PBLG)2 in DMSO at

40 °C for 24 h. The polymer products were purified by repetitive ether precipitation, dialysed against

deionized water (MWCO 7 kDa), and collected by freeze-drying. Then, the neutralized DOX was

conjugated with the hydrazide groups of mPEG-poly(Glu-Hyd) or mPEG-poly(Glu-Hyd)2 in DMSO

with a 1:1 molar ratio at 30 °C for 2 days. The products were purified by extensive dialysis against

methanol for 2 days and then dialysis against deionized water for 1 day to remove the unbound DOX.

The final product was collected by freeze-drying.

3.5. Material Characterization

Proton nuclear magnetic resonance (1H-NMR) measurements were carried out on all samples using a

Bruker DMX 600 spectrometer at 600M (Billerica, MA, USA). Fourier transform infrared spectroscopy

(FT-IR) spectra were recorded on a Bruker Tensor 27 spectrometer using potassium bromide (KBr)

disks. Polydispersity indices (PDI) of copolymers were determined by a gel permeation chromatography

(GPC) system performed on a Waters 1515 GPC instrument (Waters Corp, Milford, MA) equipped with

three Styragel columns (Waters Corp; 105, 104, and 103 Å) in tandem and a 2414 differential refractive

index detector. DMF was used as the eluent at a flow rate of 1.0 mL/min. The column temperature was

set at 35 °C. The critical aggregation concentration (CAC) of the drug-conjugated copolymers was

measured using a microplate reader (Bio Rad 680, Bio Rad, Hercules, CA, USA) with pyrene as probe [17]

by recording excitation spectra from 300 nm to 350 nm at fixed emission wavelength (390 nm). The

excitation of DOX was excluded by using DOX solutions as blanket, and the differences of fluorescent

intensity between DOX solution and drug-conjugated copolymer solutions at various concentrations

were used to calculate CACs.

Molecules 2014, 19 11928

3.6. Preparation and Characterization of the pH-Sensitive Polymeric Micelles

DOX-conjugated micelles were prepared by first dissolving the polymers in DMSO, and then

dialyzing in a dialysis tube (MWCO: 14,000) against phosphate buffer saline (pH 7.4, 10 mM) for 24 h

and deionized water for 4 h to eliminate salts. The prepared micelle solution was passed through a 0.45 μm

filter to remove any precipitate. The particle size and distribution of the micelles were determined by

dynamic laser light scattering (DLS) with a NICOMPTM 380 Submicron Particle Sizer (Particle Sizing

Systems, Santa Barbara, CA, USA) at room temperature. A laser beam at a wavelength of 632.8 nm was

used, and the scattering angle was fixed at 90° when measurements were conducted. DOX loading

content was determined using ultraviolet-visible spectrophotometry (UV-Vis, absorbance at 485 nm, T6

PERSEE, Beijing, China) by first cleaving DOX from the polymer backbone by 0.1 N HCl. The micelle

morphology was examined by transmission electron microscopy (TEM H-600, Hitachi, Tokyo, Japan)

with an accelerating voltage of 100 kV [41].

3.7. In Vitro Release Studies

The pH-responsive DOX release behavior of the micelles was investigated by dialysis diffusion [7].

Briefly, all micelles conjugating DOX were dispersed in 3 mL of medium (phosphate buffer (10 mM,

pH 7.4) or acetate buffer (10 mM, pH 5.0)) and transferred to a group of three pre-swelled dialysis tubes

(cellulose ester, MWCO: 14,000). Subsequently, the tubes were immersed in 20 mL of the same medium

and placed in a horizontal water bath shaker maintained at 37 ± 0.5 °C with a shaking speed of 100 rpm.

Then, 3 mL of release medium was replaced with fresh medium at predetermined intervals, and the

samples taken were analyzed by UV-spectrophotometry at 485 nm to determine the DOX release. All

data were obtained in triplicate using parallel experiments and were expressed as mean values.

3.8. In Vitro Cytotoxicity Assay

Human cervical carcinoma (HeLa) cells (provided by the Cell Center of the Chinese Medical

University) were cultured in RPMI 1640 medium, supplemented with 10% (v/v) fetal bovine serum

(FBS) and 1% (v/v) penicillin-streptomycin (100 U/mL penicillin and 100 μg/mL streptomycin) at 37 °C

in a humidified atmosphere with 5% CO2. The cells were harvested with 0.25% trypsin and rinsed. The

resulting cell suspensions were used in the subsequent experiments.

The cytotoxicity of the empty polymeric micelles, free DOX, and DOX-conjugated micelles against

HeLa cells was evaluated using an MTT (3-(4,5-dimethylthiazol-2-yl)-2,5-diphenyltetrazolium

bromide) cell proliferation assay[42]. Briefly, HeLa cells were seeded at a density of 3,000 cells per well

in 96-well plates for the incubation and attachment of the cells. After 24 h, the growth medium was

removed and replaced with 100 μL fresh medium containing serial dilutions of the samples (n = 5). Cells

were then incubated at 37 °C, 5% CO2, for 72 h. After predetermined periods of incubation, MTT

solution in PBS (20 μL, 5 mg/mL) was added to the sample-treated cells and incubated for the next 4 h.

The culture medium was replaced with 100 μL DMSO, followed by shaking for 10 min to dissolve the

blue formazan crystals formed. The optical density (OD) of each well was then measured at 490 nm

using a Multiscan Spectrum microplate reader (Thermo Fisher Scientific, Vantaa, Finland). Cells

Molecules 2014, 19 11929

cultured without micelle solution were used as the blank control. Cell viability was calculated according

to following equation [43]:

%100groupcontrolblankofvalueOD

grouptestofvalueOD(%) ×=viabilityCell (1)

3.9. Confocal Analysis

HeLa cells were seeded in a 6-well plate and incubated for attachment. After seeding for 24 h, the

medium was aspirated, and fresh medium containing free DOX or DOX-loading micelles was added into

the wells (DOX concentration: 10 μg/mL). After incubation for 4 h, cells were washed with

phosphate-buffered saline (pH 7.4, 0.1 M) twice for 2 min each. Next, cells were fixed with 4%

formaldehyde, washed with phosphate-buffered saline. Then, the cells were stained with Hoechst 33258

for 20 min and washed with PBS twice for 2 min each. These cells were observed under a confocal laser

scanning microscope (Zeiss LSM 710 META, Jena, Germany) with 488nm argon laser excitation.

4. Conclusions

In summary, linear and Y-shaped copolymers with DOX conjugated by an ester-amide exchange

(EAE) aminolysis reaction were synthesized and characterized. Micelles with core/shell structures were

prepared by self-assembling of these copolymers in aqueous milieu. For YMs and LMs with similar

DLC, smaller diameters were observed in YMs due to the introduction of another drug-loading chain

that shortened the lengths of the hydrophobic blocks. Also, a shorter drug-loading chain in the Y-shaped

copolymers could facilitate drug release due to more loose PEG coverage and less chain entanglement

that further achieves prompt drug release under acidic conditions. YMs based on the newly synthesized

copolymers showed a higher cytotoxicity than that of the traditional LMs, while micelles composed of

non drug-conjugated polymers showed almost no tumor-suppressive effect against HeLa cells up to a

concentration of 50 μg/mL. These results demonstrate that poly aminoacid-based Y-shaped

copolymers can be used as effective building blocks in the construction of effective drug delivery

systems and further in vivo studies are warranted.

Supplementary Materials

Supplementary materials can be accessed at: http://www.mdpi.com/1420-3049/19/8/11915/s1.

Acknowledgments

This research was supported by the National High-tech R&D Program of China 863 Program

(2012AA020305). David B Jack is gratefully thanked for correcting English in the manuscript.

Author Contributions

Conceived and designed the experiments: BS HX XT XJ. Performed the experiments: BS JJ JG.

Analyzed the data: BS JL YZ. Contributed reagents/materials/analysis tools: HZ JX. Wrote the paper:

BS XJ.

Molecules 2014, 19 11930

Conflicts of Interest

The authors declare no conflict of interest

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